Materials which are initially liquid and mouldable but that can set rapidly in situ giving immediate structural support and adhesion to surrounding tissues are of great value in bone tissue-engineering applications as well as dental, maxillofacial and orthopaedic surgeries. As they set from liquid to solid, micromechanical bonds are formed with the surrounding surfaces. The adhesive effect is especially strong with rough surrounding surfaces.
Setting of such materials may, for example, be initiated by chemical initiators or by exposure to visible or UV light (especially in cases of chemical polymerisation and cross-linking, such as in double bond containing (eg. methacrylate) polymeric formulations), or may be a result of other chemical reactions upon mixing of two components (eg acid/base in glass ionomer dental and brushite-forming bone cements) or solvent removal or evaporation from the initial liquid formulation.
For example, injectable methacrylate based dental restorative composites and adhesives and poly(methyl methacrylate) (PMMA) bone cements have been widely used for applications such tooth restoration and for fixing of orthopaedic implants. After injection of the initially fluid formulation (containing various methacrylate monomers and inorganic particles or PMMA powder in combination with liquid methyl methacrylate monomer), curing occurs, due to the presence of chemical initiators, and results in a solid material. Antibiotics or other antibacterial agents may be incorporated into dental composites or PMMA bone cements to decrease the risk of infection. Release of these agents is likely to decrease with time as it is controlled by diffusion which may be enhanced by water sorption
At present, however, there are functional limitations with all commercialised bone repair and tooth restoration products. The PMMA cements and dental restoratives discussed above are strong, but curing of large volumes generates excessive heat and material shrinkage which may cause necrosis of surrounding tissue or debonding. Additionally, if setting is slow, release of potentially toxic monomers is a problem [20-21]. PMMA also causes potential long-term biocompatibility problems, as it does not degrade in the body. [22]
With tooth restoration the composite is preferably permanent as natural repair is limited.
Since bone can regenerate, materials for bone repair should, however, if possible slowly degrade to components that may be used for tissue renewal or safely eliminated. The materials can also potentially be used simultaneously as small or large (eg DNA or protein) drug molecule controlled delivery reservoirs. If the material erodes at a constant rate from the surface then it may be possible to have linear (as opposed to declining) release of the drug at a rate commensurate with the device erosion.
Biodegradable orthopaedic fixation devices have been fabricated from various polyesters including poly lactide, glycolide or caprolactone. Polylactide screws have been shown to be useful alternatives to metal screws and implants [23].
Polylactides and polyesters are not generally injectable (although by raising the temperature above their glass transition temperature some formulations can become sufficiently fluid for moulding into a large cavity). Using polyethers as catalysts for ring-opening polymerisation of lactides, however, it is possible to produce fluid, relatively short chain poly(ether-co-ester)s. Attachment of acrylate or methacrylate end-groups then produces monomers which can cross-link and set with light or chemical cure activation. [Refs 24,25] Other injectable crosslinkable materials, including polyanhydrides [26] and polypropylene fumarates [Ref 27] have also been produced.
A problem with these materials is that controlling the rate of degradation (and hence concomitant drug release, where applicable) and mechanical properties, whilst maintaining rapid controllable set, is difficult to achieve [4-9].
A further problem, especially with more hydrophilic polyesters, can be that degradation is catalysed in the core of the material, leading to sudden catastrophic degradation instead of steady controlled surface degradation [10]. Surface degradation can be achieved through raising polymer hydrophobicity or polymer hydrolytic lability, but the need to maximise linear degradation can limit feasible polymer structures and thereby other material properties.
One method of improving control over mechanical and many other properties of polymers is through the addition of inorganic particles. In dental composites for example inorganic particles, such as silica glass, are added to methacrylate polymers to improve control over mechanical properties. The interface between the polymer and inorganic component is, however, often a point of material weakness. To overcome this problem the fillers are generally bound to the polymeric matrix phase via surface silane coupling agents but this interface may be weakened by water sorption catalysed hydrolysis of the silane [11].
As the filler loading is raised mechanical properties can improve but the viscosity of the fluids increases until, above the wet point of the filler, the formulation becomes too dry and crumbles. Smaller particles generally reduce wear and improve mechanical properties [23] but as the particles become smaller than about 5 μm, the maximum possible filler loadings (or filler/formulation wet points) can substantially decline, due to increased repulsions at the particle/matrix interface.
Filler addition to methacrylates is also known to reduce heat and shrinkage of polymerisation and to raise modulus, but can reduce the light-activated polymerisation rate in thicker samples, due to scattering effects [32, 33].
An alternative to polymeric adhesives and fillers is the use of calcium phosphate cements (CPCs). These are generally considered to be more biocompatible than the polymers and are widely used e.g. in craniofacial surgery and dental applications [28, 29]. For example, cements that form of hydroxyapatite (HA, Ca10(PO4)6(OH2)— the primary mineral component of bone, enamel and dentine) have been developed. One example involves reaction between tetracalcium phosphate and anhydrous dicalcium phosphate2CaHPO4+2Ca4(PO4)2O→Ca10(PO4)6(OH)2 
Upon mixing these phosphates with water, hydroxyapatite can slowly form. As the product crystallizes, it takes on a putty-like consistency and can be implanted or injected and contoured to a defect. The cement then completes the process and hardens, typically within ten to fifteen minutes, securing its position within the defect.
Faster setting aqueous calcium phosphate cements have also been developed using mixtures of monocalcium phosphate monohydrate (MCPM) and tricalcium phosphate (β-TCP). These two phosphates combine rapidly when mixed with water to form lower density dicalcium phosphate dihydrate (DCPD, also known as brushite) according to the expression (Refs 18-19).β-Ca3(PO4)2+Ca(H2PO4)2.H2O+7H2O→4CaHPO4.2H2O   (1)
In these cements the MCPM particles dissolve in water, then re-precipitate, solidifying the cement and forming brushite (DCPD) or monetite (dicalcium phosphate anhydrous (DCPA)) crystallites. In the body these may be slowly transformed to hydroxyapatite required for reminearalisation of bone. An excess of water is required in these cements to provide initial fluidity, and so the final materials have significant porosity thereby limiting mechanical properties and applications.
Degradation of calcium phosphate cements releases calcium and phosphate ions, which may be needed by the body to grow new bone tissue. However, disadvantages include brittleness, lack of strength and slowness of set. Although acting as effective adhesives and fillers, they generally do not provide significant support. They are also not generally useful as prolonged drug delivery devices, since drugs incorporated in the cement can be released too quickly because of the high porosity [30].
Combinations of polymers and calcium phosphates have been studied previously [44-48]. For example, various calcium and phosphate-containing particles, including phosphate based glasses, hydroxyapatite and tricalcium phosphate, have been added to degradable polymers (Refs 12-17). However, high filler loading can be restricted and, without strong interaction/bonding between the matrix polymer and inorganic phases, the interface can be a point of material weakness.
Numerous studies have shown that the addition of phosphate based inorganic particles to conventional polyesters can modify degradation and mechanical properties and also buffer acidic degradation products. However hydrophilic particle addition can be detrimental to dimensional stability, due to excessive water sorption-induced swelling.
Hydroxyapatite has previously been shown to increase the modulus of some degradable polymer composites, but the effect is rather limited (40). Additionally, hydroxyapatite formed by high temperature routes is of low solubility at pH 7 and so would be very slow to degrade (if at all) in the body. Large hydroxyapatite crystals would therefore be unlikely to provide the calcium and phosphate ions required for bone tissue regeneration.
Some acidic polymers (eg. polyacrylic acid) have also previously been added to calcium phosphate cements in an attempt to improve compressive strength, drug release characteristics and durability, but water remains the initial main continuous phase of the cement and is required in excess to give sufficient working time of the fluids. This limits the mechanical properties that may be achievable. [34,35] In addition these cements will still not have the rapid and controllable set possible with light curable methacrylates. With slow setting especially hydrophilic cements it is difficult to prevent drug “dumping” in the body before full set is achieved.